Dual-Pulsation Bi-Ventricular Assist Device

ABSTRACT

A ventricular assist device is disclosed which comprises a sac for wrapping around a portion of a heart, the sac having one or more inflatable chambers for compressing the heart when the chambers being inflated and a blood outlet made to an aorta, the blood outlet being the sole opening in the human blood path in the vicinity of heart, wherein during a systolic phase the inflatable chambers inflate while blood flows out of the aorta through the blood outlet, and during a diastolic phase the inflatable chambers deflate while blood flows into the aorta through the blood outlet.

CROSS REFERENCE

This application is a national stage of PCT/US2007/001906 filed on 25Jan. 2007, which claims priority from U.S. Patent Application No.60/763,143 filed on 30 Jan. 2006.

FIELD OF THE INVENTION

The present invention is related to ventricular assist devices (VAD),and in particular to a dual-pulsation bi-vertricular assist device(DPbi-VAD).

BACKGROUND OF THE INVENTION

In U.S. heart failure is a major public health problem whose managementconsumes about 1% of the national health care resources. Approximately 3to 4 million Americans were afflicted by heart failure, with 400,000 newcases being diagnosed each year. Heart transplantation has been the mosteffective therapy as compared to other medical treatments. Nevertheless,cardiac transplantation remained limited by the complications oflong-term immunosuppressive therapy, allograft Coronary artery diseases,and most critically, the serious shortage of donors. The annual numberof donor heart remains much constantly around 2,000. However, thepatients who are qualified to receive donor heart are estimated to be16,500 annually.

Mechanical circulation support (MCS) systems, both total artificialheart (TAH) and ventricular assist device (VAD) have been intensivelystudied, hoping to replace the role of heart transplantation forend-stage heart failure patients. Left ventricular assist device (LVAD)is versatile: in providing heart failure patients with therapiesincluding bridge to transplantation, bridge to recovery and alternativeto transplantation. The large-scale REMATCH (Randomized Evaluation ofMechanical Assistance, for the Treatment of Congestive Heart Failure)trial, which involved 19 medical centers and 129 end-stage heart failurepatients, indicates that, for patients treated by LVAD or pharmacologictherapy, one-year survival rate of the LVAD group doubles that of thepharmacologic group. Moreover, LVAD group enjoys much better quality oflife during the period of support. It has been an accepted fact incardiology society that a total implantable long-term (3-5 years)mechanical circulation support device, in particular the LVAD, willsolve the current dilemma of donor shortage for the heart failurepatients.

LVADs have been developed in recent years into a new medical modalitythat is expected to either work as a short-termbridge-to-transplantation support or replace in a long-term manner as analternative of heart transplantation. Continuous flow LVADs are smallerbut thrombogenic, and the non-pulsatile circulation support may inducemany complications in micro-circulation in end organs when usedchronically. Pulsatile LVADs are more physiologically compatible, butthe bulkiness and larger energy consumption prevents them from beingwidely adopted.

An ideal mechanical circulation support (MGS) design should possess, butnot limited to, the following characteristics: it 1) provides sufficientand adaptive cardiac support according to various physiologicalconditions or therapeutic requirements, 2) avoids blood trauma anddevice-induced complications, 3) requires simple implantation procedureand post-operational care, and 4) guarantees safety operation and allowsemergent salvage including necessary device maintenance/repair orreplacement. To date, none of the leading-edge LVAD products can meetall these requirements.

SUMMARY OF THE INVENTION

The embodiments provide a ventricular assist device (VAD), which avoidsthe major drawbacks associated with the traditional VADs.

The embodiments also provide a dual-pulsation bi-ventricular assistdevice (DPbi-VAD), which comprises a dual-pulsation mechanism enforcingco-pulsation and counter-pulsation circulation support.

The embodiments also provide a method for treating a patient suffering acardiac disease by using the dual-pulsation bi-ventricular assist deviceof the present invention.

The embodiments-achieve the DPbi-VAD by two approaches. First approach,a sac is used to directly compress the heart, especially to onlycompress the left and right ventricles. Second approach, a manifold anda blood pump is used to the flood flow out the aorta and flow in theaorta in sequence. Therefore, by properly control the timing ofcompression and pumping, the heart failure is improved.

Moreover, the two approaches could improve the heat failureindependently, in other word, the embodiments provide a ventricularassist device that is based on a sac for wrapping around a heart, andthe embodiments also provide another ventricular assist device, that isbased on a manifold for accessing blood from a human blood vessel and ablood pump for pumping flood from the human blood vessel through themanifold.

The embodiments further provide some specific designs on the sac themanifold and the blood pump.

One specific embodiment is a ventricular assist device (VAD). The VADhas a sac for wrapping around a heart and blood outlet made to an aorta.Herein, the sac has one or more inflatable chambers for compressing theheart when the chambers being inflated, and the blood outlet is the soleopening in the human blood path in the vicinity of heart. Moreover,during a systolic phase the inflatable chambers inflate while bloodflows out of the aorta through the blood outlet, and during a diastolicphase the inflatable chambers deflate while blood flows into the aortathrough the blood outlet.

The specific embodiment has some modification about the sac. Forexample, the sac could comprise a substantially rigid outer shell with ashape substantially confirm the shape of the heart the ventricularassist device is applied to. For example, the sac could comprise anelastic membrane hermetically attached to the inner surface of the outershell, wherein the space between the outer shell and the elasticmembrane forms the inflatable chamber. For example, the inflatablechambers could be placed substantially on ventricular free walls of theheart. For example, each of the inflatable chambers could have at leastone opening for de-airing thereof. For example, each of the inflatablechambers could be controllably connected to a driver, wherein theinflation of the inflatable chambers is individually adjustable. Forexample, the inflatable chambers could be inflated by a medium selectedfrom a group consisting of liquid and gas.

The specific embodiment has some modification about the blood outlet.For example, the blood outlet could comprise a manifold having a firstand a second pathway intersecting with each other at an angle, the firstpathway being completed embedded in the aorta with the second pathwayleading toward outside of the aorta. Herein, the manifold could have atleast one of the following variations: (1) the manifold could bevalveless. (2) The manifold could be made of a biocompatible materialselected from the group consisting of metal and elastic polymer. (3) Thewall thickness of the manifold could gradually decreases toward a firstand a second end of the first pathway. (4) The wall of the manifoldcould be perforated along the first pathway. (5) The wall of themanifold could be textured along the first pathway. (6) Both the firstand second pathways of the manifold could have a circular cross-section.

The specific embodiment has some modification about the blood outlet.For example, the blood outlet could comprise a blood pump having a firstcompartment connected to the blood outlet, wherein the volume of thefirst, compartment increases during the systolic phase and decreasesduring a diastolic phase. Herein, the blood pump could further comprisesa second compartment and an outer shell, the outer shell enclosing boththe first and second compartments, the volume of the outer shellremaining substantially constant during both the systolic and diastolicphases, the first and second compartment being separated by an elasticmembrane, the volume of the second compartment decreasing during thesystolic phase, and the volume of the second compartment increasingduring the diastolic phase. Herein, the blood pump could have at leastone of the following variations: (1) the second compartment could besubstantially filled with a medium, the medium being driven out of andinto the second compartment during the systolic and diastolic phase,respectively. Further, the medium is selected from the group consistingof liquid and gas. (2) Both the first and second compartments have atleast one de-airing opening.

As discussed above, the two approaches of the presented DPbi-VAD couldbe used to improve the heart failure separately. Therefore, manyembodiments could be a subset of the previous specific embodiment.

The construction and method of operation of the invention, however,together with additional objectives and advantages thereof will be bestunderstood from the following description of specific embodiments whenread in connection with the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

The drawings accompanying and forming part of this specification areincluded to depict certain aspects of the invention. A clearerconception of the invention, and of the components and operation ofsystems provided with the invention, will become more readily apparentby referring to the exemplary, and therefore non-limiting, embodimentsillustrated in the drawings, wherein like reference numbers (if theyoccur in more than one view) designate the same elements. The inventionmay: be better understood by reference to one or more of these drawingsin combination with the description presented herein.

FIG. 1 a is a schematic cross-sectional view showing a manifold andblood pump assembly of the present invention, to be implanted atascending aorta.

FIG. 1 b is a schematic view showing the manifold and blood pumpassembly in FIG. 1 a implanted at ascending aorta during pump operationat systolic phase.

FIG. 1 b is schematic view showing the manifold and blood pump assemblyin FIG. 1 a implanted at ascending aorta during pump operation atdiastolic phase.

FIG. 2 a is a schematic cross-sectional view showing a manifold andblood pump assembly of the present invention, to be implanted atdescending aorta.

FIG. 2 b is a schematic cross-sectional view showing the manifold andblood pump assembly in FIG. 2 a implanted at descending aorta duringpump operation at systolic phase.

FIG. 2 c is a schematic cross-sectional view showing the manifold andblood pump assembly in FIG. 2 a implanted at descending aorta duringpump operation at diastolic phase.

FIG. 3 a is a schematic cross-sectional view showing acompliance-matching, manifold of the present invention.

FIG. 3 b is a schematic cross-sectional view showing the hemodynamiccharacteristics of the compliance-matching manifold in FIG. 3 a duringpump operation at systolic phase.

FIG. 3 c is a schematic cross-sectional view showing the hemodynamiccharacteristics of the compliance-matching manifold in FIG. 3 a duringpump operation at diastolic phase.

FIG. 4 a is a schematic perspective view showing a stent-like manifoldof the present invention.

FIG. 4 b is a schematic cross-sectional view showing the stent-likemanifold in FIG. 4 a.

FIG. 5 a is a schematic perspective view showing a tissue-engineeredmanifold of the present invention.

FIG. 5 b is a schematic cross-sectional view showing thetissue-engineered manifold in FIG. 5 a.

FIG. 6 is a schematic cross-sectional view of a circular-shaped bloodpump of the present invention.

FIG. 7 a is a schematic view of a body-fitted blood pump of the presentinvention, in which one side of the outer shell is assumed beingtransparent.

FIG. 7 b is a schematic view of the body-fitted blood pump in FIG. 7 aat a different angle.

FIG. 8 a is a perspective view of a mixed-flow electro-hydraulic driverused iii the present invention before assembly.

FIG. 8 b is a cross-section view of the mixed-flow electro-hydraulicdriver in FIG. 8 a after assembly.

FIG. 9 a shows a flow passage in the mixed-flow electro-hydraulic driverin FIGS. 8 a and 8 b, wherein the flow passage is from blood pump tosac.

FIG. 9 b shows a flow passage in the mixed-flow electro-hydraulic driverin FIGS. 8 a and 8 b, wherein the flow passage is from sac to bloodpump.

FIG. 10 a is a perspective view of a centrifugal electro-hydraulicdriver used in the present invention before assembly.

FIG. 10 b is a partial cross-sectional view of the centrifugalelectro-hydraulic driver in FIG. 10 a after assembly.

FIG. 11 a shows a flow passage in the centrifugal electro-hydraulicdriver in FIGS. 10 a and 10 b, wherein the flow passage is from bloodpump to sac.

FIG. 11 b shows a flow passage in the centrifugal electro-hydraulicdriver in FIGS. 10 a and 10 b, wherein the flow passage is from sac toblood pump.

FIG. 12 is a schematic view showing a portable DPbi-VAD system layout ofthe present invention.

FIG. 13 is a schematic diagram showing details of the portable DPbi-VADsystem in FIG. 12.

FIG. 14 is a plot showing dual-pulsation control schedule, wherein AoP:aortic pressure, ECG: electrocardiogram, T₀, T₁, T₂, T₃: consecutiveperiods, T_(r), T_(f): blood pump filling durations, and T_(e): bloodpump ejection duration.

FIG. 15 is a plot showing the left ventricular (LV) pressure-volumerelationship vs. dual-pulsation pumping operation.

FIG. 16 a is a schematic perspective view showing a sac of the presentinvention.

FIG. 16 b is a schematic perspective view showing the sac in FIG. 16 awith a portion thereof being cut off.

DETAILED DESCRIPTION OF THE INVENTION Dual-Pulsation Design Concept

The design of an integrated co- and counter-pulsation bi-ventricularassist device proposes in the following several novel design featureswhich will improve the aforementioned inadequacies existing in thecontemporary LVAD designs. Basically, the adoption of the pulsatilepumping approach is based on two major considerations including: 1) theimplanted device functions compatibly with the human physiology; and 2)the long-term patency is guaranteed which assures both operationalsafety and quality of life for the patients. Pulsatile circulationassist is much physiologically compatible, and in order not to inducelong-term complications, in particular, in end organ micro-circulationand neurohormonal regulation, it is logically more sensible to considerpulsatile circulatory support. Owing to the fact that pulsatile devicesare non-obligatory, it is safer for patients to survive by their ownnative heart function when unexpected pump malfunction occurs. Inaddition, for pulsatile LVADs, less device-induced blood trauma andhence a reduced or no reliance on anticoagulant therapy warrants, to asignificant extent, the patient's post-operational life quality.

The contemporary switching to continuous flow design is thought, in thepresent inventors' opinion, to be a trade of exchanging mechanicalchallenges to physiological complications. Normally, a proper approachto solve the mechanical design difficulties pertaining to the LVADsshould attempt to seek solution methods in reducing the size and energyconsumption requirements, father than exchanging them withnon-physiological pumping which may very possibly induce new, unknownand refractory long-term complications to occur.

Counter-pulsation has proven to be useful for reducing systolicafterload and increasing diastolic coronary perfusion for diseasedhearts. Co-pulsation cardiac compression, however, may assist in thecardiac contraction by a direct, synchronous epicardial compression ofthe right and left ventricles. These two pumping characteristics canonly be implemented by using pulsatile devices. For eithercounter-pulsation or corpulsation circulatory support, the trigger ofmechanical actuation that imparts energy into the blood circulationrequires a precise phase control relative to the natural heart rhythm.To date, devices employing individual counter-pulsation or co-pulsationprinciple have been proposed and designed separately. A synergistic useof counter-pulsation and co-pulsation circulation supports in one singledevice forms the basis of the present bi-VAD design principle.

Imagine a mechanical device that can assist the blood circulation by acoordinated co- and counter-pulsation during heart systole, with theupstream ventricular contractility being increased while downstreamvascular afterload being reduced, the resultant effectiveness in cardiacoutput enhancement would be: much greatly elevated as compared to thoseonly support circulation via either upstream or downstream assistance.In the diastolic mode, counter-pulsation pumping enforced from withinthe aorta may assist coronary perfusion similarly as did by theconventional, well-proven intra-aortic balloon pump (IABP) devices.Moreover, during diastole another co-pulsation function of the deviceprovides a mechanical containment effect that prevents the diseasedheart from further abnormal dilation. This simultaneous use of co- andcounter-pulsations in the improvement of cardiac hemodynamics andmyocardial contraction and relaxation is termed herein, and will bereferred to, as dual-pulsation cardiac assist, and the hardware thatfacilitates this special circulation support modality is given a name as“Dual-Pulsation Bi-Ventricular Assist Device,” abbreviated DPbi-VAD.

The significance and impact of dual-pulsation cardiac assist to thepulsatile LVAD design is at least triple-fold. First, it reduces theenergy requirement of the conventional pulsatile driver design. Theretrograde energy originally wasted in extracting working fluids (air orsilicon oil) back from the pumping unit into the compliance chamber isretrieved and converted as the energy for systolic compression support.This push-and-pull type operation exerted from both up and downstreamends of the native ventricles can cut considerably down the driverenergy expenditure, resulting in a unique feature that cannot be enjoyedby any contemporary pulsatile devices. Second, the need for extra spacefor accommodating compliance chamber is eliminated. Compliance chamber,as redesigned in the present DPbi-VAD, turns an originally purelyuseless volume storage pouch into a functioning systolic compressionsupport unit. Hence, compliance chamber is no longer an indispensableliability to the pulsatile devices. The thoracic cavity spared for anatural heart to function is now dynamically shared by the compliancechamber (or the sac). Third, the inflow/outflow grafts are no longer thebypass flow passages that, must be equipped for LVAD to function. Noticethat almost all the past and contemporary ventricular assist devicesneed to use synthetic: grafts, typically those fabricated frompolyethylene terephthalate (Dacron) or expanded polytetrafluoroethylene(ePTFE), as duct passages for implementing blood drainage fromventricle/atrium and for facilitating blood re-entry, after mechanicalpressurization, back into the vasculature bed. Clinically, graftcannulation and anastomosis, which must be conducted withcardiopulmonary bypass (CPB) support, occupies a large portion of thesurgical time accompanying with substantial perioperational bleeding andthrombotic risks, it was observed statistically that manypost-operational complications such as thrombus formation, graftkinking, bleeding, and pannus overgrowth are associated with theseinflow/outflow grafts. The present single-port inflow/outflow design forfacilitating counter-pulsation circulation support revolutionizes thetraditional graft anastomosis: Virtually no conventional syntheticgrafts are required for the present DPbi-VAD and the anastomosis can becarried out by a beating heart surgery. A specially designed semi-rigidmanifold is introduced to result in a much shorter and more streamlinedpassage for blood flow to come in and out of the blood pump of DPbi-VAD.These aforementioned new characteristic features reverse: the generalimpression that pulsatile LVAD implantation is more complex anddifficult due to the necessity of implanting more and bulky componentsin a congested thoracic space.

It was found clinically that nearly 30% LVAD recipients died from rightheart failure. To date, pre-operational markers that can define thesuitability of LVAD implantation are yet satisfactorily developed.Often, LVAD recipients would require additional RVAD implantation toassist to the induced right heart failure: Otherwise, the pulmonarypressure will be elevated and the filling of the LVAD and hence thedelivered cardiac output enhancement will be impaired due to inadequateright ventricular contractility. The presently proposed DPbi-VAD solvesthis problem completely by employing a direct cardiac compression onboth right and left ventricles. With the co-pulsation operation of thesac, both pulmonary and systemic circulations are assistedsimultaneously. During this bi-ventricular support, homogeneouspressurization enforced around the free wall of the heart prevents theseptum movement which has long been an observed complication for all theLVAD implantations using left ventricular apical coring for bloodbypass. This bi-ventricular circulation support forms another importantand unique feature for the present device. When implanted with DPbi-VAD,it is believed that the risk associated with the traditional LVADimplantation will be significantly reduced owing to a balancedbi-ventricular circulation support.

The present innovative DPbi-VAD design also simplifies the implantationprocedure and greatly reduces the surgical risks. Beating-heart surgerycan be employed in the present DPbi-VAD implant procedures. For theDPbi-VAD implantation, apical coring is no longer required for unloadingthe left ventricle as well as for filling the blood pump. The inrseriesblood flow drainage and re-entry is presently facilitated by asingle-port manifold accessed, for example, from the descending aorta,making possible the intact heart VAD implantation that demands no CPBprocedure.

In summary, DPbi-VAD allows a pairing dual-pulsation synergism beaccomplished to remedy the traditional drawbacks associated with thepulsatile devices. The salutary IABP-generated counter-pulsation isretained; however, the traditional IABP shortcoming of inadequatecardiac output enhancement due to occlusive, intravascular balloondeployment is improved by using a non-occlusive, para-aortic blood pumpanastomosed end-to-side to the arterial vessel. Direct cardiaccompression offered by the sac additionally enhances the cardiaccontractility. The conversion of traditional compliance chamber intobi-ventricular-sac not only resolves the excessive space requirementproblem, but also proactively strengthens the ventricular contractilityby using otherwise wasted retrograde energy. Balanced right and leftheart support is also achieved naturally to avoid the LVAD-induced rightheart failure. This sac can further therapeutically contain the dilatedheart in the diastole mode. Graft cannulation is also revolutionized inthe present DPbi-VAD design. The mono-port inflow/outflow access toblood flow simplifies the surgical procedure and reduces the mortalityand morbidity associated with the graft-induced complications. With theincorporation of dual-pulsation assist in cardiac support, it isanticipated that the original high-energy and excessive surgical spacerequirements, as well as difficult-to-implant characteristics of thepulsatile devices will be improved significantly. In treating failinghearts, all the therapeutic hemodynamic and mechanical functions, i.e.systolic unloading, diastolic augmentation, cardiac output enhancementand passive mechanical containment are synergistically retained. Mostimportantly, DPbi-VAD is able to provide all these mechanicaltherapeutic remedies while leaving the diseased heart intact Thoseclinical and surgical advantages will inspire a new thought and protocoldevelopment for treating heart failure patients with advance MCSengagement under lower mortality and morbidity rates. Bridge-to-recoverywill be set as a primary objective for heart failure treatment, leavingbridge-to-transplantation and destination therapy the next prioritizedconsiderations.

DPbi-VAD System Description

The DPbi-VAD system could have the following six modules: 1) manifold,2) blood pump, 3) intra- and/or extra-corporeal driver systems, 4)physiological controller, 5) sac, and 6) energy/information transfersystem. Each module, is shown respectively in the following figures,with the explanations of the design features and functioncharacteristics being elucidated to illustrate its role played with andspecifications required by the whole system.

a. Manifold

Manifold constitutes a unified inflow/outflow port for DPbi-VAD toconnect to the aorta. This para-aortic manifold can be a valved orvalveless two-way duct. Each option has its own design goals to achieveand the choice of which type manifold is more suitable depends on theclinical indications, surgical complexity, and the judgment thephysician is having in mind prior to the implantation. FIGS. 1 and 2show respectively a representative manifold layout intended to beimplanted to the ascending or descending aorta. Basically the manifoldis configured by merging together two stream directions. Flow passagesare kept as streamlined as possible to reduce momentum loss as well asturbulence generation during pulsatile actuation of the blood pump.

As shown in FIG. 1 a, the ascending manifold 100A, connected to a bloodpump 200 with a quick connector 204, has a tangential distal outflowtract 101 for ejecting blood stream into the aorta 10 (FIG. 1 c) and acurved proximal duct 102 for blood pump filling (FIG. 1 b). This designaims to minimize the total pressure loss and turbulence generation,which occurs frequently in the high-speed pump ejection phase, as bloodpasses through the conduit duct. The distal end could be eithervalveless or mounted with an artificial valve 103 to regulate the bloodflow direction. A jellyfish valve is used herein because, of itslow-cost, good hemodynamic performance, low valve sound, andthrombo-resistance offered by a seamless integration to the conduitwall. Mechanical or porcine prosthetic valves can also be adopted thoughthey are more expensive. The valved manifold possesses a better pumpfilling efficacy and most importantly, it prevents the cerebral bloodretrogression during the pump filling phase. The disadvantages, ofcourse, are valve-induced hemolysis and thromboembolism, in addition tothe transvalvular energy loss due to higher flow resistance caused byvalve occlusion.

As shown in FIG. 2 a, the descending manifold 100, however, is aT-shaped conduit connected to a blood pump 200 with a quick: connector204. During the systolic phase, direct cardiac compression is offered bya sac 500 with the help from an electro-hydraulic driver 300 whileundergoing blood pump filling from the aorta 10 to the blood pump 200(FIG. 2 b) via the descending manifold 100. During the diastolic phase(FIG. 2 c), blood in the blood pump 200 is ejected into the aorta 10 viathe descending manifold 100. The intersecting angle at the junction canbe varied to yield different flow resistances hence biased perfusions tothe up and downstream blood circulation. Like the ascending manifold,the descending manifold can also be valved or valveless. Theinstallation of a prosthetic valve is helpful for a more effectivesystolic unloading provided by the blood pump 200, because this one-wayflow regulation of the descending aortic flow can avoid the volumedisplacement caused by the femoral circulation retrogression. Since thedescending aortic placement of manifold 100 and blood pump 200 involvesa longer lumen distance from the inflow/outflow port to the aortic root,the counter-pulsation actuation should take into account the phase delaydue to the finite distance traveled by the imparted pressure pulse inthe aorta before it reaches the aortic root. Descending manifoldplacement minimizes the possibility of cerebral stroke because, whenblood pump 200 ejects, the device-induced clots or microembli arerelatively difficult to convert upstream to the brain owing to largertraveling distance and smaller convection wave speed involved.

Descending valveless manifold is most attractive because it possessestwo major clinical advantages. First, descending anastomosis can beperformed with beating heart surgery. Second, artificial valve-inducedcomplications can altogether be avoided. Since a T-junction flow passageis not physiologic, therefore, vascular maladaptation that encouragesthrombus formation, intimal hyperplasia and smooth muscle cellovergrowth may occur. In order to optimize the flow and stressconditions, special designs for coping with this end-to-side anastomosisare proposed. These propositions include, but not limited to, thesubsequent design alternatives: 1) compliance-matching manifold, 2)stent-like manifold and 3) tissue engineered manifold, as described inthe following.

1. Compliance-Matching Manifold

Compliance mismatching has known to be the main causal factor forstenosis at the anastomotic junction. The jump of compliance at thegraft ends results in geometric discontinuity when subjected to bloodpressure pulsation. Geometric discontinuity causes high wall shearstress gradient being generated, at the graft/vessel junction andlow-speed recirculation flow occurring in the immediate downstream,leading to endothelial cell erosion and diffusion-induced cellproliferation around the anastomotic site. Compliance-matching graftdesign aims at eliminating this compliance discontinuity phenomenon atthe graft/vessel junctions. Shown in FIG. 3 a is the compliance-matchingmanifold design principle. The graft could be fabricated by elasticpolymers such as, but not limited to, polyurethane. Thiscompliance-matching manifold 100 has a second pathway 104 for connectingto the blood pump, and a first pathway 105 being adapted to be implantedin the aorta (or to access blood from a human blood vessel). Herein,first pathway 105 and second pathway 104 intersect with each other at anangle. The first pathway 105 possesses gradually thinning wall thicknesstoward its two ends. Since wall compliance is inversely proportional tothe product of wall thickness and Young's modulus pertaining to thegraft/vessel material, a compliance-matching graft should have asharp-edged, or zero-thickness, conduit end configuration. The graftdiameter could be designed a little bit larger (0-20%) than the diameterof the inner aortic lumen. When insertion type anastomosis is applied toplace the varying-wall-thickness graft embedded inside the aortic lumen,a tight fit of overlaid graft and aorta can be achieved using suturingmethod to result in a continuous varying compliance over the overlappedgraft/vessel region.

This manifold 100 has many hemodynamic and biological advantages. Thesmooth and compliance-matching junction design minimizes the propensityof stenosis and intimal hyperplasia occurring around the anastomoticsite. Streamlined flow pattern can be maintained as much as possible bya smooth, continuously varying composite lumen configuration. Suchlow-turbulence duct design is the optimum that an end-to-sideanastomosis can attain. Generally speaking, in the pump filling andejection phases, as shown respectively in FIGS. 3 b and 3 c, separatedflow related high-pressure stagnation zones 106 as well as low-speedrecirculation regions 107 take place in the vicinity of the T-junction.These non-physiological flow characteristics are the major factors thatcause blood cell trauma and pathologic vascular adaptations. However,for the present manifold design, and anastomosis, the liningnon-biological graft baffles the natural vessel from being afflicted bythese abnormal and mixed high-shear and low-speed flow characteristics.It is worth noticing that this baffling effect protects the surgicalwound from high pressurization, hence greatly reducing the possibilityof peri- and post-operational bleeding complications. Aside from the endregions where low-compliance dominates, the central manifold is actuallya hard-walled graft. This hard wall is more advantageous forfacilitating counter-pulsation coronary bipod perfusion. Diastolicaugmentation is found sensitive to the vessel wall elasticity. Theinsertion type anastomosis, in fact, replaces a portion of the naturalvessel by a hard-walled graft. As blood pump ejects, the higheststagnation flow region is directly contained and resisted by thehard-walled manifold. The overall wall compliance is therefore decreasedwhich is welcome as considered by the facilitation of diastolicaugmentation.

2. Stent-Like Manifold

Another embodiment of the manifold is achieved by way of stenting themanifold graft. FIGS. 4 a and 4 b show a schematic of this designconcept. The manifold 100 is constructed using biocompatible metalmaterials such as stainless steel, titanium or titanium alloys. On theaorta-embedded duct portion (the first pathway 105), holes orperforations 108 of various shapes and porosity ratios can bedistributed to form perforated walls 109, allowing endothelium cellmigration in the same manner as observed in the stented bipod vessel.The stent porosity can be varying, with higher porosity (lower modulus)regions, located around the graft/vessel junctions. Althoughcompliance-matching is relatively difficult to achieve presently ascompared to the aforementioned compliance-matching polymer graft design,the endothelialization, after certain period after implantation, willsandwich and embed the stent-like manifold inside the aortic lumen,resulting in a neon composite vessel of higher rigidity in the centralregion and a softer and smoother geometric/compliance transition aroundthe ends. For arteries of larger diameter (> 6˜9 mm), re-stennosis israrely found clinically due to the higher flow velocity effect,Thrombo-resistance is guaranteed by the neonintima layer produced by thestent characteristics of the graft structure. Lower compliance of thestented arterial portion may well enhance the diastolic augmentation asdoes the manifold.

3. Tissue-Engineered Manifold

Tissue-engineered manifold is another variant of the compliance-matchinggraft design. The graft material used can be those elastic polymersmentioned previously. Tiny micropores or cavities 110 of smalldimensions (30˜300 microns) can be manufactured and distributed aroundcertain desired inner or outer wall surface areas on the first pathway105 of the manifold (see FIGS. 5 a and 5 b). These textured polymersurfaces work as the scaffold for cells in the blood stream to adhereand grow. The anchored dense thrombi will further encourage blood cellinteractions to occur. A heterogeneous surface containing platelets,monocytes, macrophages, foreign-body giant cells, lymphocytes, etc. willbe deposited after device implantation. Over time, a neonintimaproliferated with endothelial cells will populate over the textured areaof the manifold. Preferably these tissue-engineered neonintima islocated at the graft/vessel junctions. This may further enhance thejunction performance in terms: of better compliance-matching, smoothergeometry transition, and stronger graft adherence to the artery.

b. Blood Pump

Blood pump works alternately as a reservoir to receive the blood volumeand an ejector to propel the stored blood volume back to the artery.This blood pump 200 comprises a single-port design, with an elasticmembrane 205 separating the blood in the first compartment 201 and thesilicon working fluid in the second compartment 202, as illustrated inFIG. 6. Biocompatible material (e.g. polyurethane) is used to constructthis blood pump. The basic form, or a circular-shaped geometry, helpsflow wash-out effect be fully accomplished in the pump chamber. Theinflow/outflow port 203, however, may be placed tangentially witheccentricity relative to the pump centerline to help develop flow swirl.The stroke volume of the pump can range from 30 to 100 c.c. or larger,depending on the thoracic space that is allowed and the propelling powerthe driver is designed to deliver.

FIGS. 7 a and 7 b show another variant of the blood pump design similarto that shown in FIG. 6, wherein like parts are designated by samenumerals: The shape of this blood pump 200 looks like a flat, curvedellipsoid. Salient angles are avoided to result in a streamlinedcontour. The outer shell is configured by conforming its contour to theinner thoracic cage wall. This body-Fitted design allows the maximumusage of the thoracic space which leads to a minimal interference withthe lung. The stroke volume of the pump could range from 30 to 100 c.c.or larger.

In both circular-shaped and body-fitted pump designs, blood and workingfluid are separated by a elastic membrane. For circular-shaped bloodpump shown in FIG. 6, the elastic membrane 205 has a zero-stress shapewhich is nearly identical to the pump shell configuration. The elasticmembrane 205 is attached and sealed around the central inner peripheryof the pump shell. For body-fitted blood pump shown in FIGS. 7 a and 7b, however, the elastic membrane, or actually a sac 205 with similar butsmaller shape to the outer shell, attaches to the inflow/outflow port203 tangentially. In both designs the hard outer shell restrains thestretching of the sac, and hence prolongs the service life of theelastic membrane by limiting the elastic membrane strain well below thefatigue-related threshold. As circular-shaped pump ejects the elasticmembrane 205 moves upward and barely touches the shell inner surface,leaving a small amount of blood reserve remaining in the pump. In thenext pump filling phase this residual blood volume will be washed awayand mixed with the newly replenished blood. As for body-fitted pump infull-volume ejection, the sac 205 is compressed fully until the oppositeelastic membranes thereof may touch each other with minimal residualvolume of the first compartment 201 in the sac 205 remained; and whenpump fills, the sac 205 resumes its original zero-stress, wrinkle-freefull shape, with the volume of the second compartment 202 diminishes,which is formed between the sac 205 and the outer shell of the bloodpump 200. For partial-volume ejection, however, blood stored can stillbe ejected completely out of the pump, though requiring several morestrokes. The stasis-free characteristic of the present pump designs isattributed to the vortex-washing effect provided by the shell andelastic membrane contour in conjunction with the unsteady pulsatile flowmotion.

De-airing opening is placed in the blood pump wall for air removal. Thelocation is selected around the uppermost region where air bubbles arecollected when blood pump is implanted via left thoracotomy. Forcircular-shaped pump shown in FIG. 6, the de-airing opening 206 isplaced in the vicinity of the elastic membrane attachment rim. Forbody-fitted pump, however, the de-airing opening 206 is located at thebottom of the sac where the sat attachment and fixation stub is housed,as shown in FIGS. 7 a and 7 b.

As shown in FIGS. 6, 7 a and 7 b, the upper end of the bipod pump has aquick connector 204 mounted around the throat to yield a convenientquick assembly or detachment of the pump 200 to the manifold. Silicongasket is used to enable a tight integration of the blood pump with themanifold. At the bottom end, however, the pump shell converges into acircular duct 208 to allow connection to the driver. Similar quickConnector design like that used for manifold/blood pump integration isadopted for efficient pump/driver installation. Shuttling silicon oilflowing into and out of the blood pump can be accomplished by the use ofan intra- or extra-corporeal driver system. Additional de-airing opening207 is built on the pump casing shell to aid in air removal whenassembling blood pump together with the driver.

c. Intra- and Extra-corporeal Driver Systems

1. Intra-Corporeal Driver System

The present intra-corporeal driver design adopts the hydraulic pumpingprinciple. The designed electro-hydraulic (EH) driver consists of twoelectrically-driven direct current (DC) brushless motors, one for torquegeneration and another for flow direction regulation, in additional toan impeller and a switching valve. The use of pressurized hydraulicfluid, or silicon oil, enables a power deployment with literally noconstraints posed on the position and orientation of the actuatedmechanism. The present EH driver actuates blood pump and sacasynchronously by referencing the electrocardiogram (ECG) waveform.Cannulas, through which pressurized hydraulic working fluid runs, areused for power transmission. Quick connectors enable fast and convenientassembly/detachment and orientation adjustment of the EH driver body asconnected to the blood pump and/or the sac.

For pulsatile flow pumping, the search of an isolated optimal designpoint is neither critical nor practical because during pump operation aloop, rather than a fixed point, shows on the pump performance map.Often, an overall high-efficiency plateau spanning a certain operationalloop range should be looked for. In order to ease the implantation, moreoften than not, the pursuit of a miniature LVAD design demands theanatomic or space consideration a more prioritized design criterion,leaving the high operational hydraulic efficiency the next objective toachieve.

In the present DPbi-VAD driver design, the two, DC brushless motors areimmersed in the silicon oil bath contained in the EH driver. The heatgenerated comes mainly from the amateur windings and the electroniccontrollers. The sloshing motion of the oil flow may serve as aneffective cooling mechanism for converting and redistributing thedissipated heat to the surfaces of the whole DPbi-VAD, including outershells, cannulas, and elastic membranes in contacting with thecirculating blood stream. Heat, hence, can be transferred locally byconduction to the tissues which surround the VAD implant or globally byconvection throughout the human circulation system to the entire body.There will be literally no hot spots produced in the presentelectro-hydraulic design, which is an important factor for increasingthe reliability of the electronics and thus for prolonging the servicelife of the DPbi-VAD in operation.

There are two electro-hydraulic driver designs that are presented in thefollowing. One uses mixed-flow and the other centrifugal impellerdesign, depending respectively on the emphasis of either efficiency orhead-rise that is stressed.

1.1 Mixed-Flow Driver

A mixed-flow electro-hydraulic driver constructed according to one ofthe preferred embodiments of the present invention is shown in FIGS. 8 aand 8 b, which is assembled with the flowing elements/parts:

-   301 Driver Casing,-   302 Base Cap,-   303 Bearings-   304 Lock Ring,-   305 Torque Motor Casing,-   306 Torque Motor Stator,-   307 Torque Motor Rotor,-   308 Bearing,-   309 Lock Ring,-   310 Switching Valve Body,-   311 Bearing,-   312 Switching Valve Head,-   313 Switching Valve Outflow Duct,-   314 Stepping Motor Rotor,-   315 Stepping Motor Stator,-   316 Switching Valve Inflow Duct,-   317 Bearing, and-   318 Mixed-flow Impeller.

The impeller 318 is housed in an inner switching valve formed by theswitching valve outflow duct 313 and the switching valve inflow duct 316which are non-rotatably engaged with each other, and the inner switchingvalve is rotatably disposed in the stationary casing formed by theswitching valve body 310 and the switching valve head 312. A DC brushless torque motor, formed by the motor casing 305, the motor stator 306,and the motor rotor 307, drives this mixed-flow impeller 318 with atypical speed range of 5,000-12,000 rounds per minute (RPM). Around11,000 RPM, at least 20 liter/min (LPM) flow rate should be deliveredagainst 120 mmHg pressure rise at the aortic site. The impeller wasdesigned and optimized by using computational fluid dynamics (CFD)analysis. A stepping motor, formed by the motor rotor 314 and the motorstator 315, is installed around the waist of the switching valve (313,316). This stepping motor (314, 315) can align together theinflow/outflow aperture pairs which are drilled respectively on the sidewalls of the stationary casing (310, 312) and the inner rotatingswitching valve (313, 316). The stepping motion is regulated by acontroller using ECG waveform as the reference for developingcounter-pulsation pumping. Both motors revolve uni-directionally so asto minimize the energy consumption due to direction reverse. Two flowpassage routes, as determined by the switching valve motion, will directthe pressurized oil flowing back and forth between the blood pump andthe sac via an inflow/outflow port 320 and an outflow/inflow port 319formed on the driver casing 301. As shown in FIGS. 2 b and 2 c, theblood pump 200 is connected to the driver 300 via a cannula 321, and thesac 500 is connected to the driver 300 via another cannula 322. Thisintegrated driver and flow passage system is shown in FIG. 9 a and FIG.9 b. Note that all electric and mechanical components are immersed inthe oil chamber. Excellent cooling and lubrication can be achieved tomaintain the long-term patency of this EH driver.

Centrifugal Driver

An electro-hydraulic driver using a centrifugal impeller for pumpingsilicon oil constructed according to one of the preferred embodiments ofthe present invention is shown in FIGS. 10 a and 10 b which is assembledwith the flowing elements/parts:

-   401 Driver Casing,-   402 Dome Valve,-   403 Torque Motor,-   404 Centrifugal Impeller,-   405 Dome Valve Head,-   406 Switching Connector,-   407 Bearing,-   408 Stepping Motor Rotor,-   409 Stepping Motor Stator, and-   410 Driver Head.

The flow passage which allows oil shuttling between sac and blood pumpby this centrifugal driver is illustrated in FIGS. 11 a and 11 b. Theimpeller 404 was also designed using advanced CFD package assisted by anin-house design optimization procedure. The inlet and outlet diametersof the radial impeller and the blade heights at inflow and outflowplanes are appropriately chosen to attain a high pump efficiencyspanning over a wide operational speed range. The design objective ofthe present centrifugal EH driver is set for delivering a volume flowrate (>20 LPM) against a pressure gradient of 200 mmHg at a rotationalspeed around 4000 to 8000 RPM. Metallic or ceramic ball bearing could beused for supporting the power transmission shaft which is also the rotoraxis of the DC brushless torque motor 403. Both inner and outer spinningrotor designs can be used. Since silicon oil can purge the heat anddebris generated within the bearing, a longer service life of thepresent EH driver can thus be guaranteed.

A bell-shaped switching valve, formed by the dome valve 402 and the domevalve head 405, regulates the flow and hydraulic power transmissiondirections. This bell-shaped switching valve is driven by a DC brushlessstepping motor containing the stepping motor rotor 408 and the steppingmotor stator 409. Windows are formed circumferentially around the upperend of the dome valve 402, and between the dome valve head 405 and theswitching connector 406. The gap and recess created between thebell-shaped switching valve and the driver casing 401 of the EH driverforms an annular passage for the silicon oil flow to enter the impellerinlet. The inflow, coming either from the systolic sac or out of theblood pump, circulates around the annular gap, climbs up and entersthrough the opening windows of the dome and finally gets devoured intothe impeller inlet eye.

The impeller-pressurized oil, however, is collected by a volute 412whose outer wall is the dome valve 402. An aperture 413 is opened on thevolute wall which as rotates and aligns alternately with the two sidecannulas connecting respectively to blood pump or systolic sac, fulfillsthe power deployment mission. A pair of “V-shaped” arms 414 and 415,which extrude respectively out from the left and right sides of the EHdriver casing 401, forms the inflow/outflow tracts for shuttling theworking fluids back and forth between the blood pump and the sac. EHdriver can be viewed as a pressurization unit which alternately sloshesoil flow between its two ends, one being the low-pressure ventricularside (˜0 mmHg) and the other the high-pressure (80-120 mmHg) aorticside. Preferred high-performance outflow tract angle is assigned to thevolute while pumping the blood pump. Extra loss will be incurred duringthe subsequent withdrawal of oil flow back from the blood pump due tothe sharp turn of the outflow tract at the volute exit. However, thisloss would be compensated by the higher preload condition of the aortawhen shuttling oil back from the blood pump to compress the ventricles.

2. Extra-Corporeal Driver System

Extra-corporeal driver system can be either a bedside or a portableunit. The bedside model is used primarily in the intensive care unit(ICU) in a hospital. The portable model, however, is designed foroutpatients when discharged after the DPbi-VAD implantation and beingcleared from the post operational care period.

Extra-corporeal driver systems could be propelled by either pneumatic orhybrid pneumatic/electro-hydraulic power source, as described below. Inorder to minimize the infection possibility caused by the percutaneousdrive lines, thin pneumatic lines like those used by IABP are adopted.Basically, for all the extra-corporeal systems described below, theintra-thoracic fluid power actuation lines that connect respectivelywith the implants, i.e. the blood pump and the sac, are thin pneumatictubes in which inert, low molecular weight gas such as helium isrunning. At the outlet of each intra-thoracic pneumatic drive line, adermal button is implanted under the skin, allowing convenient quickattachment/disconnection of the inner pneumatic drive line to theoutside power source system.

For all the extra-corporeal drivers the intra-thoracic powertransmission design is the same. By plugging on/off the exterior drivelines which connect either to the bedside console or to the portableDPbi-VAD driver unit, the patients implanted with the presentintra-thoracic device may be either bedridden or ambulant depending onthe medical treatment and the life style the patients are having.

2.1 Bedside Driver, System

Because the present DPbi-VAD uses the same operational principles asdoes the IABP, the IABP driver consoles Can be employed as the driversystem for pumping DPbi-VADs. Adaptor can be designed to connectDPbi-VAD to the IABP consoles, with pneumatic volume and pumpingpressure adjusted according to the DPbi-VAD operational requirements. Ingeneral, two IABP drivers with asynchronous actuations of blood pump andsac are required for driving a DPbi-VAD system. Separate application ofcounter-pulsation and co-pulsation can also be selected for patients whorequire only one circulation support modality under certain clinicalconditions.

Another embodiment of the bedside DPbi-VAD is accomplished using adesignative driver system. This bedside driver is almost an identicalequipment derived from the portable DPbi-VAD driver as illustratedbelow. The only differences lie in first, the electric power supplymodule allows for receiving the wall alternative electric current asidefrom the battery-supplied direct current, and, second, a moresophisticated monitoring/display/adjustment system is equipped for themedical personnel to reference and control.

2.2 Portable Driver System

Portable driver is designed for patients who are cleared from the ICUstage. The present portable system is a variant of the intra-corporealdriver described previously. FIGS. 12 and 13 illustrate the idea of howthis hybrid pneumatic/electro-hydraulic driver system is constructed.This system, in fact, is derived from the intra-corporeal EH driver byrelaxing the space/anatomic constraint. The impeller size and therotational speed of the torque motor can be modified to account forlosses incurred in the longer fluid pathways associated with theextra-corporeal system. The volute inflow/outflow tracts are alignedtangentially to the volute body so sharp turn in the fluid pathway iseliminated to achieve a better hydraulic efficiency. Connected to theinflow/outflow cannulas 321 and 322 of are now two reservoirs 323 and324 instead of the original blood pump 200 and sac 500 that weredesigned for the intra-corporeal system. Each reservoir is divided bylayers of elastic membranes 325 and 326 into two partitions filled bysilicon oil and helium gas, respectively. At the proximal heliumpartition side of each reservoir a pneumatic line 327 (328) exits andconnects to the intra-thoracic unit via a specially designedanti-infection skin button 329 (330). When operated, the impeller-drivenshuttling silicon oil will create compression and vacuum power onto thetwo neighboring reservoirs 323 and 324, hence driving the helium air andresulting in a simultaneous actuation of the blood pump 200 and the sac500 implanted in the patient's thoracic chest.

Multiple-redundancy can be enforced using multiple hybrid drivers.Double-redundancy using two driver units EDH1 and EDH2 arranged inparallel is exemplified in FIG. 13. As pump malfunction is detected, theelectronic controller 600 will immediately send out a command signal tostart the spared driver and a non-stopping, continuous driving motioncan thus he guaranteed. Control valves 601, 602 are installed in thejunctures of the fluid pathways to regulate the actuated fluiddirection. An alternative, intermittent operation of the two EH driverscan also be designed using control logic programmed in the controller.By interchangeably providing appropriate idle time to each EH driver, ahoverall prolonged service life of the portable driver system can beattained.

Each reservoir 323 (324) has a pneumatic line, connected to the heliumreplenishment system 700. Appropriate helium volume can be selected fortreating different heart failure syndromes encountered clinically. Thereplenishment procedure is similar to those being implemented with theIABP systems. After reconnection of the extra-corporeal pneumatic lines327 and 328 and intra-corporeal pneumatic lines 331 and 332, heliumvolume has to be completely drained first and then replenished to therequired volume.

Comparing to the contemporary pneumatic driver systems which use bulkyand heavy reciprocal engine and storage tanks, the presentturbomachinery system is much light-weighted and quieter. Subsystemssuch as battery pack 603, electronic controller 600, and the hybridpneumatic/electro-hydraulic units 800 are integrated into boxedcanisters, which could easily be carried by a patient using wearablejacket, as depicted in FIG. 12. It has always been welcome to have a VADsystem with minimal modules that have to be implanted inside the humanbody. For the present portable DPbi-VAD, the complex electro-mechanicaland power supply/control modules are all placed extra-corporeally. Notonly regular inspection and maintenance of the driver and controller canbe allowed, the heat generated by the driver during operation can alsobe dumped easily to the atmosphere. In case of emergent machinemalfunction, a quick replacement of the faulty unit can also be carriedout. Great mobility and safety operation is hence guaranteed by thisportable driver system. Since the pulsatile operation of blood pump andsac is non-obligatory, stoppage of the DPbi-VAD is allowable for acertain period of time (a few minutes to a fraction of hour per se).This will further provide convenience and freedom for patients to puton/off the DPbi-VAD jacket for routine events, such as changing clothesand/or taking bath that are frequently happening in a real life.

d. Physiological Controller

The objective of the present controller design aims at developing ascheduled timing and forcing level control for the EH driver. Asuccessful enforcement of systolic unloading, diastolic augmentation andepicardial compression calls for a delicate phase manipulation inrelation to the heart rhythm. Systolic unloading and epicardialcompression are enforced as a pair of a single actuation and theinitiation is best set at the time when the ventricles are just about toContract. Diastolic augmentation, however, is initiated during heartdiastole, when the aortic valve is just closed (starting from thedicrotic notch of the aortic pressure trace) and the coronary arterialwalls begin to relax. All these actuation controls need to use ECG oraortic pressure waveform as a base of reference. Suppose that ECG signalis adopted for pumping control, algorithm has to be developed torecognize the R-wave out of other wave characteristics on the ECG signaltrace. Similar waveform recognition method can readily be developed ifaortic pressure is selected as the sensor signal.

Cardiac output can be taken as a product of stroke volume multiplied byheart rate. Physiological heart regulation is realized by adjustingstroke volume and heart rate autonomously via nervous and hormonalcontrol. Usually higher heart rate corresponds to larger heart musclecontractility and hence elevated stroke volume. Therefore, aphysiological controller that mimics the natural cardiac regulation canbe built based on heart rate alone, leaving EGG signal the only requiredcontrol input. Since R-wave is tracked, the pumping frequency of the EHdriver can be determined, which may be identical, or in proportional tothe detected heart rate. This characteristic makes the present deviceoperate in response to the physiological circulation need. It should benoted that as much frequent electro-hydraulic pumping ejection isfulfilled by a faster switching valve aperture crossing, the cross-flowresistance of the aperture will increase accordingly. Theimpeller-delivered hydraulic pressure should therefore be elevated toovercome the aperture loss and the higher ejection inertia required. Thespeed control of the stepping and torque motors should therefore becoordinated using appropriate control logic, resulting in a situationthat circulation assistance is physiologically enforced to meet thecardiac output requirement and the therapeutic purposes as well. In caseof arrhythmia, a default pumping scenario of fixed pumping frequencywill dictate irrespective of the arrhythmic ECG signal.

The actuator system of the present controller consists of two motors,driving respectively the switching (or dome) valve for frequency orheart rate control and the impeller for mechanical power deliveryregulation. Because mixed-flow and centrifugal EH drivers are similar inthe sense of controller development, we shall illustrate in thefollowing the control design of the centrifugal pump only. Equivalentcontrol design for mixed-flow pump can be achieved by a minor adjustmentof the parameters involved.

For the centrifugal EH driver, a stepping motor is mounted onto the domevalve head for position and motion control. In controlling the oil flowdirections the stepping motor revolves continuously of intermittentlyand uni-directionally. In each heart beat the pumping control can bedivided into two phases. The first is the systolic unloading control inthe aortic side and, spontaneously, the epicardiac compression in theventricular side. The second is the diastolic augmentation control whilethe sac functions like a slaved oil pouch. For both systolic and,diastolic actuations, each pumping control basically consists of threesteps, starting with the initiation control and then followed by aposition and a duration control. The initiation control decides when thecontrol action commences. The position control drives the dome apertureto align with the inflow/outflow ports. The duration control, however,determines how long the pump ejection should last upon the completion ofthe aperture alignment. All these six control steps are accomplished inone heart beat cycle using the prior detected time period as a base fordefining the control algorithm. FIG. 14 illustrates the control schedulein conjunction with the hemodynamics, valve motion and cardiogram inConsecutive cardiac cycles. These timing and period of controlactuations are further translated into the pressure-volume relationshipof the left ventricle, as shown in FIG. 15.

On FIG. 14 there are four consecutive heart cycles that are shown toillustrate the present control algorithm. The default position is theposition where impeller volute aperture aligns with the inflowport/cannula leading to the sac. This is also the position the referencezero-degree angle is defined and reset after each revolution. At cycle 0the control is yet initiated and the switching valve is located at thedefault zero-degree position. The driver sits still with pumpingdirection oriented for systolic unloading. As two consecutive R-wavesare detected, the cycle period T₀ is calculated and two time intervalsT_(f) and T_(e), which are proportional to T₀, are determinedaccordingly. A pair of parameters, which define the proportions of T_(f)and T_(e) relative to the available cycle period T₀, is preset initiallyin the control logic. Diastolic augmentation commences at the end of thefirst elapsed time interval T_(f). At this instant, position controlwould quickly move and align the dome valve aperture with the outflowport/cannula to the blood pump. Diastolic augmentation is then enforceduntil the T_(e) interval is ended. Upon completion of diastolicaugmentation, the stepping motor is actuated to continue the rotationand reposition the valve aperture in alignment with the inflowport/cannula to the sac. Systolic unloading is immediately activatedupon the completion of aperture alignment. Roughly speaking, a residualtime duration will take up the remaining pumping cycle before the nextR-wave onset. This means that the aortic pressure has already beendeclined prior to the heart systole. When new R-wave is detected, anupdated cycle period T₁ becomes available and the following two timeintervals T_(f) and T_(e) are re-calculated. This cyclic control commandsequence will recursively be created and executed with new R-waves beingconsecutively detected.

A photo detector sensor pair is mounted on the dome valve and the drivercasing to help reposition the default position of the dome valve as one360-degree revolution is completed in a cycle. The zero-degree referenceposition is redefined in each cycle as the photo detector pair alignedtogether. In the present design the zero-angle position is set where theapertures are aligned to enforce systolic unloading. This repositioncontrol can prevent accumulated angular drift and assures in each beatthe volute aperture Would align correctly with the outflow tract.

Owing to the fact that larger through-flow aperture corresponds to lowerejection loss, the control logic should be selected, under the powerlimit of the stepping motor, to maximize as fast as possible therotational acceleration for positioning the dome valve aperture inalignment with the right or left outflow tracts. Systole/diastole ratiocan be adjusted to yield an optimal diastolic augmentation and theassistance to the systemic circulation as well. In situation ofarrhythmia, or heart rate larger than certain limit (i.e. 100 beats perminute), the controller will ignore the cardiogram and gives either afixed rate or a reduced 1-to-2 or 1-to-3 pulsation command to yield anoptimal cardiac support, same as those supplied by the IABP controllogic.

A DC brush less torque motor provides the main driving force for theimpeller. The power is supplied by a battery pack and regulated by aspeed controller. This torque motor is expected to deliver a poweroutput ranging from 20 to 30 watts. As heart rate increases, theaperture resistance and the arterial pressure both elevate, requiring alarger pumping force to achieve the desired cardiac output requirement.Either passive or active torque motor control can be adopted to copewith this physiological demand. A predetermined piece-wise RPM targetschedule, for example can be set for a heart beat range for the passivecontrol with an autonomous RPM tracking control being implemented in themotor drive. Sophisticated active physiological control can also beconsidered, which calls for more sensors or information be supplied toreflect the demand of the body or the regulation rule the cardiacfunction is supposed to possess.

e. Sac

Integration of bi-ventricular systolic compression withcounter-pulsation circulation assist is a unique feature of the presentinvention. Originally, for pulsatile LVAD devices, the need of acompliance chamber has been a necessary liability because pulsationrequires additional volume to undergo the back-and-forth displacement ofthe stroke volume. Hence, the space required for device implantation isdoubled. The implementation of sac turns this disadvantage intoadvantage. When sac inflates, the ventricles contract, and vice versawhen sac deflates. The space for accommodating this, sac, therefore,does not occupy too much additional space since it dynamically sharesthe space with the natural ventricles.

According to the preferred embodiments the present sac consists of twosheets of polyurethane (PU) foils. One, the outer shell, which issemi-rigid but non-distensible, or is substantially rigid, is embeddedwith textile nets to restrain its stretching deformation. The other,being an elastic diaphragm used to assist heart Contraction, is a thinflexing membrane made by solution dip method. The size and morphology ofthese two PU foils are, in principle, chosen to be almost identical tothose of the heart in the end-of-diastolic condition. For dilatedfailing hearts, appropriate sizing of the sac can be employed to limitfurther pathological dilatation. Clinically it was observed that heartwith pathological dilatation may have very irregular shape. To achievean effective epicardial compression, the fitness of sac in regard to thediseased heart shape is very important. Customized sac can be fabricatedusing the CT-scanned images taken prior to the implantation.

Depicted in FIGS. 16 a and 16 b is a schematic of the present sac 500.The outer shell 501 and the elastic membrane 502 are fused togetheraround the sac rim by the PU solvent or other methods to result in asealed volume closure. In other words, the elastic membrane 502hermetically attached, to the inner surface of the outer shell 501.Moreover, the combination of the outer shell 501 and the elasticmembrane 502 form one or more inflatable chambers. A conduit 503 isprovided on the non-flexing outer shell 501, which allows silicon oil beinjected into and/or withdrawn from the sac 500. A opening for de-airing504 is installed oh the outer shell to help facilitate air removalwhenever deemed necessary. In general, each inflatable chamber has oneopening for de-airing. It is worthy noticing that the elastic membrane502 has no attachment spots with the outer shell except the peripheralrim. This free-movement characteristic of the elastic membrane providesthe sac a greater ability in shape adaptation. Moreover, the presentelastic membrane design is particularly meaningful in minimizing therelative movement occurring between the elastic membrane and the heartskin. Myocardial contusion or scar tissue formation can thus be greatlyavoided. Further, each of the inflatable chambers could be controllablyconnected to a driver, wherein the inflation of the inflatable chambersis individually adjustable.

In the installation of the sac, the pericardium is first dissected andthe fluid drained to allow the insertion of the sac. Sac is theninserted through the pericardial opening to allow a snug wrapping of thesac around the heart. Note that the sac should only wrap around thefight and left ventricles. This sizing control can prevent the deliveredhydraulic pressure from compressing the atria. This consideration isessential for maintaining a reasonable and low end-diastolic-pressure,avoiding the impairment of the subsequent filling of the ventricles.Sometimes, compressing atria during ventricle systole might causepremature opening of the atrio-ventricular valves, which is particularlyundesirable for patients with incompetent valves. In all, care must beexercised in contouring the sac configuration, making atria cleared outfrom the wrapping of the sac. Isolation of atria from systoliccompression support is necessary; otherwise, the venous return andpulmonary vein pressures will be elevated due to the applied compressionforce.

The thin PU sheets in contact with the ventricles are very compliant andshape conformal. The remaining myocardial fluid on the heart skin worksas a liquid film to help compel out the air trapped initially when sacis mounted on the ventricles. It was found in the in-vivo experiments,that typically a few pumping strokes will squeeze out air bubblesinitially trapped, making the inner membrane of the sac firmly attachedon the ventricle skin. Coronary perfusion would not be significantlyimpaired because the sac works synchronously with the heart. Indiastole, both heart and sac are relaxed, which encourages coronaryperfusion. Since the sac membrane is firmly attached on the heart skin,during diastolic augmentation stroke the vacuum created in the sac maygenerate a suction effect on the ventricle free wall, which isbeneficial for increasing the end-of-diastolic volume and hence thestroke volume. In addition, this suction effect helps to, expand thecoronary arterial lumen also, creating another salutary effect inenhancing the coronary flow.

Fixation of sac holds an important role on upholding the performance ofdirect cardiac compression. As ventricle systoles, the resultantreaction force exerting on the sac points approximately in the apicaldirection in the septum plane. This reactant force is mainly resisted bythe oil conduit strut which is supported in turn by the implanted EHdriver. A cuff ring 505 is made around the conduit 503 located at thebase of the sac 500. This base cuff ring 505 can be tightly sutured ontothe epicardium to help control the orientation and fixation of the sac.If necessary, additional fixation can be made by sewing or anchoringother sac cuffs (not shown in the figures) onto the surrounding pleuraltissues and/or bone structures.

Bi-ventricular circulation support is provided by the sac since bothleft and right ventricles are wrapped in. In the systolic unloadingmode, a volume of silicon oil no greater than the cardiac stroke volumewill be sloshed into the sac. This shuttled volume, typically 30˜90 c.c,will be redistributed in the inward volume recess created by thecontraction of both left and right hearts. Note that this shuttledvolume occupies only a portion of the total stroke volume contributed bythe left and right heart contraction. The effective epicardialcompression action, therefore, only takes place in the initialcompression phase. In other words, the ventricles are boosted in theforemost part starting from, or just before, isovolumctric contractionto the early stage of blood ejection. For the rest contraction period,the sac will be passively carried by and moved inward together with thecontracting heart muscle. This limited inward power stroke of the saccan prevent the fight heart from being overtly compressed. Nevertheless,direct contact of the heart muscle by the sac still effectively helpsthe initial muscular shortening which is known to be crucial in reducingthe myocardial oxygen consumption.

In summary, the functions of the sac are multitude. First, it works asan artificial pericardial pouch which protects the heart from contusionand absorbs the vibration or impact momentum exerted from outside.Second, in diastole, it plays the role of restraining heart from furtherabnormal dilation. Third, and most importantly, it assists both rightand left heart contractions to result in a balanced circulation support,avoiding the frequently observed LVAD-induced complication of rightheart failure due to left heart support. Direct sac compression canincrease the contractility of the diseased ventricles and, asincorporated with counter-pulsation circulation assist, not only helpfill the blood pump but also prevent excessive unloading which is knownto be detrimental to coronary perfusion augmentation.

f. Energy/Information Transfer System

Both percutaneous and transcutaneous energy/information transfer systemscan be considered. For percutaneous system, the implanted components arelesser, which eases the surgical operation at the price of increasingthe possibility of post-operational infection. Electrical energy andsensor/commanding signal transfer is more reliable and efficient as hardwires are used. Interferences caused by wireless, energy/datatransmission can be minimized. Besides, the intra-thoracic heat releasewill also be reduced because many heat generating electronic equipments,such as motor controllers, battery, and data acquisition and processingunits, are not intra-corporeally implanted.

Transcutaneous transfer system is welcome because it provides thehighest quality of life for the patients. Because extra components, suchas internal charge battery set and electronic controllers must beimplanted, the bulkiness of the implantable VAD system increases. Inorder to transfer the internally produced heat out from the implants,the heat generating components had better be packaged in the EH driver,which will enlarge the EH driver volume size accordingly.

The features as described above which are not claimed in the pendingclaims would be claimed in divisional applications of the presentapplication.

1. A ventricular assist device comprising: a sac for wrapping around aportion of a heart, the sac having one or more inflatable chambers forcompressing the heart when the chambers being inflated; and a bloodoutlet made to an aorta, the blood outlet being the sole opening in thehuman blood path in the vicinity of heart, wherein during a systolicphase the inflatable chambers inflate while blood flows out of the aortathrough the blood outlet, and during a diastolic phase the inflatablechambers deflate while blood flows into the aorta through the bloodoutlet.
 2. The ventricular assist device of claim 1, wherein the saccomprises an outer shell being able to substantially maintain a contoursubstantially confirm to a contour of the heart the ventricular assistdevice is applied to.
 3. The ventricular assist device of claim 2,wherein the sac further comprises an elastic diaphragm having a contoursimilar to the contour of the outer shell, the elastic diaphragm beingattached to the rim of the inner surface of the outer shell and coveringthe inflatable chambers.
 4. The ventricular assist device of claim 1,wherein the inflatable chambers are placed substantially on ventricularfree walls of the heart.
 5. The ventricular assist device of claim 1,wherein the inflatable chambers are inflated by a medium selected from agroup consisting of a liquid and a gas.
 6. The ventricular assist deviceof claim 1, wherein the blood outlet comprises a manifold having a firstand a second pathway intersecting with each other at an angle, the firstpathway being completed embedded in the aorta with the second pathwayleading toward outside of the aorta.
 7. The ventricular assist device ofclaim 6, wherein the manifold is valveless.
 8. The ventricular assistdevice of claim 6, wherein the manifold is made of a biocompatiblematerial selected from the group consisting of metal and elasticpolymer.
 9. The ventricular assist device of claim 6, wherein the wallthickness of the manifold gradually decreases toward a first and asecond end of the first pathway.
 10. The ventricular assist device ofclaim 6, wherein the wall of the manifold is perforated along the firstpathway.
 11. The ventricular assist device of claim 6, wherein the wallof the manifold is textured along the first pathway.
 12. The ventricularassist device of claim 6, wherein both the first and second pathways ofthe manifold have a circular cross-section.
 13. The ventricular assistdevice of claim 1 further comprising a blood pump having a firstcompartment connected to the blood outlet, wherein the volume of thefirst compartment increases during the systolic phase and decreasesduring a diastolic phase.
 14. The ventricular assist device of claim 13,wherein the blood pump further comprises a second compartment and anouter shell, the outer shell enclosing both the first and secondcompartments, the volume of the outer shell remaining substantiallyconstant during both the systolic and diastolic phases, the first andsecond compartment being separated by an elastic membrane, the volume ofthe second compartment decreasing during the systolic phase, and thevolume of the second compartment increasing during the diastolic phase.15. The ventricular assist device of claim 14, wherein the secondcompartment is substantially filled with a medium, the medium beingdriven out of and into the second compartment during the systolic anddiastolic phase, respectively.
 16. The ventricular assist device ofclaim 15, wherein the medium is selected from the group consisting of aliquid and a gas.
 17. The ventricular assist device of claim 16, whereinboth the first and second compartments have at least one de-airingopening.
 18. A ventricular assist device comprising: a sac for wrappingaround a portion of a heart, the sac having one or more inflatablechambers for compressing the heart when the chambers being inflated; anda manifold having a first and a second pathway intersecting with eachother at an angle, the first pathway being completed embedded in theaorta with the second pathway leading toward outside of the aorta,wherein during a systolic phase the inflatable chambers inflate whileblood flows out of the aorta through the second pathway of the manifold,and during a diastolic phase the inflatable chambers deflate while bloodflows into the aorta through the second pathway of the manifold.
 19. Theventricular assist device of claim 18, wherein the sac comprises anouter shell being able to substantially maintain a contour substantiallyconfirm to a contour of the heart the ventricular assist device isapplied to.
 20. The ventricular assist device of claim 19, wherein thesac further comprises an elastic diaphragm having a contour similar tothe contour of the outer shell, the elastic diaphragm being attached tothe rim of the inner surface of the outer shell and covering theinflatable chambers.
 21. The ventricular assist device of claim 18,wherein the inflatable chambers are placed substantially on ventricularfree walls of the heart.
 22. The ventricular assist device of claim 18,wherein the wall thickness of the manifold gradually decreases toward afirst and a second end of the first pathway.
 23. The ventricular assistdevice of claim 18, wherein the wall of the manifold is perforated alongthe first pathway.
 24. The ventricular assist device of claim 18,wherein the wall of the manifold is textured along the first pathway.25. The ventricular assist device of claim 18 further comprising a bloodpump having a first compartment connected to the second pathway of themanifold, wherein the volume of the first compartment increases duringthe systolic phase and decreases during a diastolic phase.
 26. Theventricular assist device of claim 25, wherein the blood pump furthercomprises a second compartment and an outer shell, the outer shellenclosing both the first and second compartments, the volume of theouter shell remaining substantially constant during both the systolicand diastolic phases, the first and second compartment being separatedby an elastic membrane, the volume of the second compartment decreasingduring the systolic phase, and the volume of the second compartmentincreasing during the diastolic phase.
 27. The ventricular assist deviceof claim 26, wherein both the first and second compartments have atleast one de-airing opening.